Typically, in computed tomography (CT) imaging systems, a rotatable gantry includes an x-ray tube, detector, data acquisition system (DAS), and other components that rotate about a patient table that is positioned at the approximate rotational center of the gantry. X-rays emit from the x-ray tube, are attenuated by the patient, and are received at the detector. The detector typically includes a photodiode-scintillator array of pixelated elements that convert the attenuated x-rays into visible light photons within the scintillator, and then to electrical signals within the photodiode. The electrical signals are digitized and then received and processed within the DAS. The processed signals are transmitted via a slipring (from the rotational side to the stationary side) to a computer for image reconstruction, where an image is formed.
The gantry typically includes a pre-patient collimator that defines or shapes the x-ray beam emitted from the x-ray tube. X-rays passing through the patient can cause x-ray scatter to occur, which can cause image artifacts. Thus, x-ray detectors typically include an anti-scatter grid (ASG) for collimating x-rays received at the detector.
Third generation multi-slices CT scanners typically include detectors having scintillator/photodiodes arrays. These detectors are positioned in an arc where the focal spot is the center of the corresponding circle. These detectors generally have scintillation crystal/photodiode arrays, where the scintillation crystal absorbs x-rays and converts the absorbed energy into visible light. A photodiode is used to convert the light to an electric current. The reading is typically linear to the total energy absorbed in the scintillator.
Typically, CT systems obtain raw data and then reconstruct images using various known pre-processing and post-processing steps to generate a final reconstructed image. That is, CT systems may be calibrated to account for x-ray source spectral properties, detector response, and other features, to include temperature. Raw x-ray data are pre-processed using known steps that include offset correction, reference normalization, and air calibration steps, as examples.
In recent years, the development of volumetric or cone-beam CT technology has led to an increase in the number of slices used in CT detectors for computed tomography systems. The detector technology used in large coverage CT enables greater coverage in patient scanning by increasing the area exposed, by using back-illuminated photodiodes. The increase of the number of slices results in an increase in the width of the detector in Z-axis (e.g., along the patient length). Because it is impractical to build very large modules in monolithic structure to cover 160 mm or more in the Z-axis, due to manufacturing cost and reliability concerns, in one example stack smaller modules (mini-modules) are built along the Z-axis.
However, image quality in a CT canner is dependent on several components in the system such as the detector, the x-ray tube and high voltage generator, the system and component geometry, and the thermal management, etc. In third generation CT scanners, the detector for example, typically has very strict specifications to ensure good image quality and some of these requirements include but are not limited to: a) stability of the detector over time and temperature, b) focal spot drift, c) stable and high light output over the lifetime of the detector, etc.
One important factor is related to thermal management. Typically, detectors are calibrated at a known temperature and imaging data is obtained. However, a variety of factors can cause the detectors to fall out of calibration. For instance, temperature within the room or suite (in which the CT scanner is placed) can vary or drift throughout the day, or power dissipation within the CT scanner can cause temperature changes depending on how heavy the scanner is used. In fact, there are numerous factors that can lead to variation in detector temperatures, which can lead to the detector falling out of calibration.
And, the components within the detector itself can be sensitive to temperature, and the gain of the detector components may decrease with increased temperature. To counter this sensitivity, in one example it is known that detectors or detector assemblies may be heated during calibration and use to ensure a stable temperature during operation. However, heating the detector components may have the deleterious or counter effect of reducing overall dose efficiency, given that the gain may be reduced. Thus, on one hand stable operation and extended periods of operation may be achieved by heating the detectors or detector assemblies. However, such heating may come at the cost of dose efficiency. Further, as detector assemblies have grown along the Z-axis in recent years, the task of thermal management may be more difficult as components within the detector may be prone to thermal drift.
Thus, there is a need to improve thermal management within a CT detector or CT detector module.